JP7240415B2 - System and method for ultrasound screening - Google Patents

Description

The present invention relates to the field of ultrasound, particularly fetal ultrasound screening.

Due to the spread of assisted reproductive technology accompanying the rise in maternal age, the number of twin births has increased significantly. Due to the high risk of twin pregnancies, regular monitoring of fetal well-being is important. Fetal health is typically assessed by measuring the fetal heart rate using a Doppler ultrasound transducer. Twin pregnancies require manual placement of two independent ultrasound transducers in the maternal abdomen, which requires skill and experience.

Clinical problems experienced during twin fetal heart rate measurements may be caused by both hearts lying in the same measuring range of one transducer. The measured Doppler signals contain information related to both fetal hearts and the fetal heart rate estimation algorithm fails to extract correct fetal heart rate information for either fetus.

Furthermore, both transducers can be misdirected at the same fetal heart and measure the same heart rate, thereby missing the heart rate information of the other fetus. This can lead to the potential problem that one of the fetuses will go unnoticed entirely. The position of the heart may also change over time due to fetal motion, which may require repositioning of the ultrasound transducer. This introduces additional uncertainty into the measured fetal heart rate.

Additionally, it is possible that the heart rates coincide by chance even though both ultrasound transducers are correctly positioned. In this case, the system may not be able to distinguish whether both measured heart rates belong to the same fetus or are from two different Doppler sources. Furthermore, registration of the maternal heart rate may occur rather than the fetal heart rate(s). This is because the pulsating maternal artery is located within the measurement volume.

What is needed, therefore, is a system that more reliably acquires heartbeat information from multiple fetuses without requiring significant additional hardware.

The invention is defined by the claims.

According to an embodiment according to one aspect of the invention, an ultrasound system is provided, the system comprising:
An ultrasound transducer array, said ultrasound transducer array having a plurality of transducer elements configured to conform to a body of a subject, wherein at least two ultrasound transducer elements of said plurality of transducer elements are located in a region of interest an ultrasound transducer array that acquires a plurality of ultrasound signals from the region of interest at different orientations relative to
A processor for receiving ultrasonic signals acquired by the ultrasonic transducer array,
splitting a plurality of ultrasound signals based on signal depth;
Compute the Doppler power for each ultrasound signal split,
determining, for each ultrasound signal, the depth of the fetal heartbeat based on the Doppler power of the split of each ultrasound signal;
a processor for determining a fetal heart region based on the determined fetal heart rate and the positions of the at least two ultrasound transducers.

This system allows the fetal heartbeat to be located more accurately, especially when there are multiple fetal hearts (twins) or when maternal blood vessels are captured within the ultrasound signal.

Doppler power reflects the total motion within a given measurement volume, or sample volume. The signal depth used to split the ultrasound signal determines the depth of this sample volume. The shape of the transmitted ultrasound beam determines the length and width of the sample volume. Using the known position (orientation) of the transducer, the position of the sample volume within the imaging area may be determined.

Typically, the sample volume with the strongest Doppler power indicates the location of the fetal heart. The 3D visualization further clarifies the position within the imaging region and any further Doppler sources (eg maternal blood vessels).

Using arrays configured to easily fit the shape of the subject clinically eliminates the complex task of precisely positioning two ultrasound probes to acquire two fetal heartbeats in the case of twins. reduce or eliminate the need for a physician

In an embodiment, the calculation of Doppler power comprises:
calculating the Doppler signal based on the splitting of the ultrasound signal; and calculating the root mean square value of the Doppler signal over a predetermined time period, thereby calculating the Doppler power.

In a further embodiment, the time period is 1 second or more, such as 2 seconds or more.

Thus, the risk of missing heartbeats within the measurement window is reduced or eliminated.

In one configuration, determining the depth of the fetal heartbeat comprises comparing Doppler power to a threshold power.

Thus, power values below a given value are discarded, thereby removing small motions due to factors other than fetal heartbeat.

In a further configuration, determining the depth of the fetal heartbeat further comprises clustering Doppler power values equal to or greater than the threshold power.

Thus, the Doppler power due to the fetal heartbeat is grouped, which provides an indication of the fetal heart region.

In a further configuration, clustering is performed by a two-element Gaussian mixture model or a k-means clustering model.

In one embodiment, the processor is further configured to calculate a median fetal heart rate based on the fetal heart rate of the identified fetal heart beats.

In an embodiment, the system further includes a display, the display configured to display the fetal heart region to the user.

In a further embodiment, the processor is further configured to determine the location of each ultrasound signal split within the region of interest, and the display displays a fetal heart region in association with each ultrasound signal split. further configured.

Displaying the fetal heart region in relation to the entire region of interest simplifies identification for the user. Furthermore, it can provide a specific indication of the wrong fetal heart region, for example, when a maternal blood vessel is identified as the fetal heartbeat.

In one configuration, the ultrasonic transducer array further includes a sensor and the processor is further configured to determine the curvature of the ultrasonic transducer array based on the output of the sensor.

By determining the curvature of the array of transducer elements, the position and orientation of the ultrasound transducer with respect to the region of interest can be more accurately known.

In a further configuration, the sensor comprises:
strain gauge,
Acceleration sensor,
It has one or more of a piezoelectric sensor and a camera.

In embodiments, the plurality of ultrasonic transducers comprises:
piezoelectric transducers; and CMUTs.

According to an example according to one aspect of the invention, there is provided an ultrasound imaging method, the method comprising:
acquiring a plurality of ultrasound signals with at least two ultrasound transducers having different orientations with respect to the region of interest;
splitting the plurality of ultrasound signals based on signal depth;
calculating the Doppler power for each ultrasound signal split;
determining, for each ultrasound signal, the depth of the fetal heartbeat based on the Doppler power of each ultrasound signal split;
and determining a fetal heart region based on the determined fetal heartbeat and the positions of the at least two ultrasound transducers.

In one arrangement, determining the depth of the fetal heartbeat comprises:
comparing the Doppler power to a threshold power;
and clustering Doppler power values greater than or equal to the threshold power.

According to an example according to one aspect of the invention there is provided a computer program comprising computer program code means, which computer program is configured to implement the method when said computer program is run on a computer. .

These and other aspects of the invention will be apparent from the embodiments described below and will be described with reference to these embodiments.

1 is a diagram of an ultrasound diagnostic imaging system to illustrate general operation; FIG. 1 shows a schematic representation of an ultrasound system according to the invention; FIG. Fig. 3 shows a schematic representation of the division of a received echo signal; Fig. 3 shows a further schematic representation of an ultrasound system according to the invention; 3D visualization of a sample volume of a measurement region containing two fetal hearts; FIG. Fig. 3 illustrates the method of the invention;

For a better understanding of the invention and to show more clearly how it can be realized, reference is made, by way of example only, to the accompanying drawings.

The invention will be described with reference to the figures.

The detailed description and specific examples, while indicating exemplary embodiments of apparatus, systems and methods, are intended for purposes of illustration only and are intended to limit the scope of the invention. It should be understood that the These and other features, aspects, and advantages of the apparatus, systems, and methods of the present invention will become better understood from the following description, the appended claims, and the accompanying drawings. It should be understood that the drawings are schematic representations only and are not drawn to scale. It is also understood that the same reference numbers are used throughout the drawings to designate the same or similar parts.

The present invention provides an ultrasound system including an ultrasound transducer array and a processor. An ultrasound transducer array has a plurality of transducer elements configured to conform to a subject's body. Further, at least two ultrasound transducer elements of the plurality of ultrasound transducer elements are configured to acquire multiple ultrasound signals from the region of interest in different directions with respect to the region of interest. The processor is configured to receive ultrasound signals acquired by the ultrasound transducer array. The processor is further configured to split the plurality of ultrasound signals based on the depth of the signal and calculate Doppler power for each ultrasound signal split. For each ultrasound signal, the processor determines the depth of the fetal heartbeat based on the Doppler power of the split of each ultrasound signal; based on the determined fetal heartbeat and the positions of the at least two ultrasound transducers; Identify the fetal heart region.

General operation of an exemplary ultrasound system will first be described with reference to FIG. 1, with an emphasis on the signal processing capabilities of the system. This is because the present invention relates to the processing of signals measured by transducer arrays.

The system has an array transducer probe 4 with a transducer array 6 for transmitting ultrasound waves and receiving echo information. The transducer array 6 comprises CMUT transducers, piezoelectric transducers made of materials such as PZT or PVDF, or any other suitable transducer technology. In this example, transducer array 6 is a two-dimensional array of transducers 8 capable of scanning either a 2D plane or a three-dimensional volume of the region of interest. In another example, the transducer array may be a 1D array.

The transducer array 6 is coupled to a microbeamformer 12 that controls reception of signals by the transducer elements. Microbeamformers are transducers, as described in U.S. Pat. (commonly referred to as "groups" or "patches") can be at least partially beamformed.

Note that the microbeamformer is completely optional. Additionally, the system includes a transmit/receive (T/R) switch 16, to which the microbeamformer 12 can be coupled, switching the array between transmit and receive modes, and when the microbeamformer is not used. , to protect the main beamformer 20 from high energy transmit signals when the transducer array is directly manipulated by the main system beamformer. Transmission of ultrasound beams from the transducer array 6 is directed by a transducer controller 18 . This controller is coupled to the microbeamformer by a T/R switch 16 and to the main transmit beamformer (not shown). This controller may receive input from the user's manipulation of the user interface or control panel 38 .

Controller 18 may include transmit circuitry configured to drive the transducer elements of array 6 (directly or via a microbeamformer) during a transmit mode.

In a typical line-by-line imaging sequence, the beamforming system within the probe may operate as follows. During transmission, the beamformer (which may be a microbeamformer or a main system beamformer in some embodiments) activates the transducer array or sub-apertures of the transducer array. A sub-aperture may be a one-dimensional line of transducers or a two-dimensional patch of transducers within a larger array. In transmit mode, the focusing and steering of the ultrasound beams produced by the array or subapertures of the array are controlled as described below.

Upon receiving the echo signals backscattered from the target, the received signals undergo receive beamforming (described below) to align the received signals, and if subapertures are used, the subapertures are It is then shifted, for example by one transducer element. A shifted sub-aperture is then activated and the process repeated until all transducer elements of the transducer array are activated.

For each line (or sub-aperture), the sum of the received signals used to form the associated line of the final ultrasound image is the voltage signal measured by the transducer elements of the given sub-aperture during the reception period. is the sum of The line signals obtained following the beamforming process below are typically referred to as radio frequency (RF) data. Each line signal (RF data set) produced by the various sub-apertures then undergoes additional processing to produce the lines of the final ultrasound image. Variations in the amplitude of the line signal with time contribute to variations in the brightness of the ultrasound image with depth, where high-amplitude peaks correspond to bright pixels (or groups of pixels) in the final image. Peaks appearing near the beginning of the line signal represent echoes from shallow structures, and peaks appearing later in the line signal represent echoes from structures at greater depths within the object.

One of the functions controlled by transducer controller 18 is the direction in which the beam is steered and focused. The beams may be steered straight from the transducer array (perpendicular to the transducer array) or steered at different angles for a wider field of view. Steering and focusing of the transmit beam may be controlled as a function of the activation time of the transducer elements.

In general ultrasound data acquisition, two methods can be distinguished: plane wave imaging and "beam steered" imaging. The two methods are distinguished by the presence of beamforming in transmit mode (“beam steered” imaging) and/or in receive mode (plane wave imaging and “beam steered” imaging).

First for the focusing function, by activating all transducer elements simultaneously, the transducer array produces a plane wave that diverges as it passes through the object. In this case the beam of ultrasound remains unfocused. Introducing a position-dependent time delay in the activation of the transducer allows the wavefront of the beam to converge to a desired point, called the focal zone. The focal zone is defined as the point where the lateral beamwidth is less than or equal to half the transmit beamwidth. Thus, the lateral resolution of the final ultrasound image is improved.

For example, if the time delay causes the transducer elements to activate serially, starting from the outermost element of the transducer array and ending at the center element of the transducer array, then along the center element, a given distance away from the probe forms a focal zone. The distance of the focal zone from the probe varies based on the time delay between each subsequent round of transducer element activation. After the beam passes through the focal zone, the beam begins to diverge, forming a far field imaging region. Note that in the focal zone located close to the transducer array, the ultrasound beam diverges quickly in the far field, resulting in beamwidth artifacts in the final image. Generally, the near field located between the transducer array and the focal zone shows little detail due to the large overlap of the ultrasound beams. Thus, changing the position of the focal zone can lead to large changes in final image quality.

Note that in transmit mode only one focus can be defined unless the ultrasound image is divided into multiple focus zones (each of which may have a different transmit focus).

Additionally, upon receiving echo signals from within the object, it is also possible to perform reception focusing by performing the reverse processing of the above-described processing. In other words, the input signal may undergo an electronic time delay before being received by the transducer elements and passed to the system for signal processing. The simplest example of this is called delay-and-sum beamforming. It is possible to dynamically adjust the receive focusing of the transducer array as a function of time.

Turning now to the function of beam steering, the correct application of time delays to the transducer elements can impart a desired angle to the ultrasound beam exiting the transducer array. For example, by activating the transducers on a first side of the transducer array and then activating the remaining transducers in a sequence ending on the opposite side of the array, the wavefront of the beam is angled toward the second side. be done. The magnitude of the steering angle relative to the transducer array normal depends on the magnitude of the time delay between subsequent transducer element activations.

Furthermore, it is possible to focus the steered beam. Here the total time delay applied to each transducer element is the sum of both the focus time delay and the steering time delay. In this case the transducer array is called a phased array.

For CMUT transducers that require a DC bias voltage for their activation, the transducer controller 18 can be coupled to control a DC bias control 45 for the transducer array. DC bias control 45 sets the DC bias voltage applied to the CMUT transducer elements.

For each transducer element of the transducer array, analog ultrasound signals, typically called channel data, enter the system via receive channels. In the receive channel, partially beamformed signals are generated from the channel data by the micro-beamformer 12 and then passed to the main receive beamformer 20, where the beamformed signals are partially beamformed from the individual patches of the transducer. The resulting signals are combined into a fully beamformed signal called radio frequency (RF) data. The beamforming performed at each stage may be performed as described above or may include additional functions. For example, the main beamformer 20 may have 128 channels, each receiving partially beamformed signals from patches of tens or hundreds of transducer elements. Thus, signals received by thousands of transducers of a transducer array can efficiently contribute to a single beamformed signal.

The beamformed received signals are coupled to a signal processor 22 . Signal processor 22 can process the received echo signals in a variety of ways, including bandpass filtering; decimation; I and Q component separation; It functions to separate linear and non-linear signals so as to allow identification of non-linear (higher harmonics of the fundamental frequency) echo signals returned from . The signal processor may perform additional signal enhancements such as speckle reduction, signal combining, and noise reduction. The bandpass filter in the signal processor may be a tracking filter whose passband slides from a higher frequency band to a lower frequency band as echo signals are received from greater depths, thereby Noise at higher frequencies from greater depths, which typically lack anatomical information, is rejected.

The beamformer for transmission and the beamformer for reception are implemented in different hardware and can have different functions. Of course, the receive beamformer is designed with the properties of the transmit beamformer in mind. In FIG. 1 only the receive beamformers 12, 20 are shown for simplicity. In a complete system there is a transmit chain with a transmit microbeamformer and a main transmit beamformer.

The function of the microbeamformer 12 is to provide an initial combination of signals to reduce the number of analog signal paths. This is typically done in the analog domain.

Final beamforming is done in the main beamformer 20, typically after digitization.

The transmit and receive channels use the same transducer array 6 with fixed frequency bands. However, the bandwidth occupied by a transmit pulse can vary depending on the transmit beamforming used. The receive channel can capture the entire bandwidth of the transducer (which is the classical approach), or use bandpass processing to capture the desired information (e.g., harmonics of the dominant harmonic). Only the bandwidth that contains can be extracted.

The RF signal may then be coupled to a B-mode (ie, luminance mode, or 2D imaging mode) processor 26 and a Doppler processor 28 . The B-mode processor 26 performs amplitude detection on the received ultrasound signals to image internal structures such as organ tissue and blood vessels. For line-by-line imaging, each line (beam) is represented by an associated RF signal, the amplitude of which is used to generate the luminance values assigned to the pixels of the B-mode image. The exact location of a pixel within an image is determined by the location of the associated amplitude measurement along the RF signal and the line (beam) number of the RF signal. B-mode images of such structures are formed in the harmonic or fundamental image modes described in US Pat. No. 6,283,919 (Roundhill et al.) and US Pat. No. 6,458,083 (Jago et al.). or a combination of both. A Doppler processor 28 processes the temporally differentiated signals resulting from tissue motion and blood flow for the detection of moving objects such as blood cell flow within the image field. The Doppler processor 28 typically includes a wall filter with parameters set to pass or reject echoes returned from selected types of material within the body.

The structural and motion signals produced by the B-mode and Doppler processors are coupled to scan converter 32 and multiplanar reformatter 44 . Scan converter 32 places the echo signals into the desired image format in the spatial relationship in which they are received. In other words, the scan converter functions to convert the RF data from a cylindrical coordinate system to a Cartesian coordinate system suitable for displaying ultrasound images on the image display 40 . For B-mode imaging, pixel brightness at a given coordinate is proportional to the amplitude of the RF signal received from that location. For example, the scan converter may place the echo signals into a two-dimensional (2D) sector format, or into a pyramidal three-dimensional (3D) image. The scan converter can overlay the B-mode structural image with colors corresponding to motion at points within the image field. Here the Doppler estimated velocity produces a given color. Combining the B-mode structural and color Doppler images, tissue motion and blood flow within the structural image field are depicted. A multi-planar reformatter converts echoes received from points within a common plane of a body volumetric region into an ultrasound image of that plane, as described in U.S. Pat. No. 6,443,896 (Detmer). do. Volume tenderer 42 transforms the echo signals of the 3D data set into projected 3D images viewed from a given reference point, as described in US Pat. No. 6,530,885 (Entrekin et al.). .

A 2D or 3D image is transferred from the scan converter 32, the multiplanar reformatter 44, and the volume tenderer 42 to the image processor for further enhancement, buffering, and temporary storage for display on the image display 40. 30. The image processor may be configured to remove certain imaging artifacts from the final ultrasound image. This artifact is e.g. acoustic shadowing caused by strong attenuators or refraction, e.g. posterior enhancement caused by weak attenuators, e.g. reverberation artifacts when highly reflective tissue interfaces are placed in close proximity, and so on. Additionally, the image processor may be configured to process certain speckle reduction functions to improve the contrast of the final ultrasound image.

In addition to being used for imaging, blood flow values generated by Doppler processor 28 and tissue structure information generated by B-mode processor 26 are coupled to quantification processor 34 . The quantification processor produces structural measurements such as organ size and gestational age, as well as measurements of various flow conditions such as blood flow volume fraction. The quantification processor may receive input from the user control panel 38, such as points in the image anatomy at which measurements are to be made.

Output data from the quantification processor is coupled to graphics processor 36 for reproducing measurement graphics and numbers along with images on display 40 and outputting audio from display device (image display) 40 . The graphics processor 36 can also generate graphic overlays for display with the ultrasound images. These graphic overlays can include standard identifying information such as patient name, image date and time, imaging parameters, and the like. For these purposes, the graphics processor receives input from the user interface (or panel) 38, such as patient name. A user interface is also coupled to the transmit controller 18 for controlling the generation of ultrasound signals from the transducer array 6 and thus the generation of images generated by the transducer array and ultrasound system. The transmit control function of controller 18 is only one of the functions performed. The controller 18 also takes into account the mode of operation (given by the user), the corresponding required transmitter configuration, and the bandpass configuration in the analog-to-digital converter of the receiver. The controller 18 may be a state machine with fixed states.

The user interface is also coupled to a multi-planar reformatter 44 for plane selection and control of multiple multi-planar reformatted (MPR) images. This can be used to perform quantified measurements in the image field of an MPR image.

FIG. 2 shows a schematic diagram of an ultrasound system 100 having an ultrasound transducer array 110 and a processor 120, which may be, for example, one or more of the processors 26, 28, 30, 34 described above.

The ultrasound transducer array 110 has a plurality of transducer elements 130 and is configured to conform to a subject's body 140 . At least two ultrasound transducer elements of the plurality of ultrasound transducer elements are configured to acquire ultrasound signals from the region of interest 150 at different orientations with respect to the region of interest. Each individual transducer element 130 is configured to transmit and receive ultrasound waves. The transducer elements can comprise piezoelectric transducers or CMUT cells.

Transducer array 110 may be configured to conform to subject's body 140 in several ways. For example, multiple transducer elements may be embedded in a flexible silicone layer.

In other words, the transducer array may be configured to conform to the subject's body to ensure that the transducer elements have good contact with the body surface. Additionally, the material layer placed under the element between the transducer element and the target may be selected to have an appropriate acoustic impedance suitable for propagation of ultrasound waves. The transducer array may be made from any suitable material, for example by integrating the transducer elements into a cloth or belt, which may be wrapped around the subject's body.

Additionally, flexible arrays need not be completely closed. For example, individual elements may be interconnected by any flexible connector piece that defines the approximate position of the elements relative to each other.

Alternatively, individual transducer elements 130 can be attached directly to the subject's skin in a manner similar to ECG measurement electrodes that attach directly to the skin.

In addition, a subset of the elements (eg, a 7 element transducer subarray) may be placed on a rigid plate, which may be placed on the skin. These multiple sub-arrays may be used to cover a large area while following the curvature of the object being measured.

In the example shown in FIG. 2, ultrasound system 100 is employed to measure the heart rate of a fetus. More specifically, the transducer array 110 is positioned adjacent the maternal abdomen to irradiate the fetal region.

Due to the fact that the transducer array 100 is flexible and positioned on the maternal abdomen, each transducer element is directed at a particular angle to the maternal abdomen. Knowing the approximate curvature of the array allows estimating the position of each sample volume that makes up the 3D ultrasound image captured with the ultrasound system. The approximate curvature of the transducer array can be estimated from the average curvature of the pregnant mother's abdomen.

Alternatively, the ultrasound transducer array 110 may further include a sensor 160 and the processor 120 is configured to determine the curvature of the ultrasound transducer array based on the output of the sensor. The sensors can include one or more of strain gauges, accelerometers, piezoelectric sensors, and cameras. For example, a camera may be used to determine the curvature of the array and may also be used to determine the position of the array on the maternal abdomen.

The known positions of the individual transducer elements within the 2D array and the curvature of this array make it possible to obtain the position of the transducer elements relative to the patient, thereby providing the x-position and y-position of the heart's position within the measurement volume. - Position is estimated.

Processor 120 is configured to receive ultrasound echo signals acquired by the ultrasound transducer array. Upon receiving the signal, the processor is configured to split the signal based on the depth of the signal. This signal division is further described below with reference to FIG.

FIG. 3 shows schematically the splitting of the received echo signal. A first axis 170 represents the displacement z between the above-described transducer array 110 positioned adjacent to the maternal abdomen 180 and the fetal heart 190 .

A second axis 200 shows an exemplary ultrasound signal 210 that may be produced by an ultrasound transducer of the transducer array. The second axis corresponds to the first axis proportional to the velocity of the ultrasound signal in tissue.

A third axis 220 shows received ultrasound echo signals 230 reflected by the fetal heart 190 . A similar signal is received by each active transducer element of the transducer array. The received signal is divided by time gate 240 .

Dividing the received echo signals with multiple time gates 240 allows estimating the depth (z-position) of the heart location within the measurement volume.

Multiple Doppler signals are calculated from different depths for each transducer element. In other words, the Doppler power is calculated for each ultrasound signal split. This can be done by setting multiple time gates, also called range gates, during the Doppler power calculation process. For each transmitted ultrasound burst, one sample n of the received ultrasound echo signal yig(n) is acquired, as shown in FIG. 3, where i represents the element index and g the range gate index. A common Doppler processing scheme can be used for each range gate. The total number of acquired Doppler signals is therefore the product i*g.

Calculating the Doppler power may include calculating the Doppler power by calculating a Doppler signal based on the division of the ultrasound signal and calculating a root mean square value of the Doppler signal over a predetermined time period. .

In other words, the Doppler signal is calculated for each division of the received echo signal 230 . The length of this calculated Doppler signal depends on the size of the range gate. A root-mean-square value of the Doppler signal is calculated over a range gate that defines a predetermined time period. For example, the time period is 1 second or more, such as 2 seconds or more.

Stated another way, in the example, the power of each Doppler signal is calculated using the root mean square value over a two second time period to ensure that the heart rate is always measured. The Doppler power Pi,g reflects the total motion within a particular sample volume. The range gate index determines the depth of the sample volume and the shape of the transmitted ultrasound beam determines the width of the sample volume.

Thus, by noting the range gate index of the detected motion, the depth of the fetal heartbeat can be determined based on the Doppler power of each ultrasound signal split.

Determination of fetal heartbeat depth may include comparing Doppler power to a threshold power. That is, by identifying Doppler powers above a given threshold, background motion of the imaging region, eg, the maternal abdomen, may be discounted. Thus, fetal heart beats may be isolated within one or more range gates, thereby identifying the sample volume in which the fetal heart is located.

Determining fetal heartbeat depth may further include clustering Doppler power values equal to or greater than the threshold power. Automated clustering provides spatial separation of the Doppler sources within the measurement volume, thereby providing automatic detection of fetal heart rate and/or position. In one example, clustering may be performed using a two-element Gaussian mixture model.

After thresholding the Doppler power values, all sample volumes that are good candidate volumes for measuring fetal heart rate are identified, and given the estimated curvature of the transducer array, the positions of the sample volumes are identified. can be A Gaussian mixture model assumes that the data (positions in the sample volume) come from a Gaussian distribution. In the example of measuring twin heart rates, the data are fitted to a two-factor Gaussian model because it is known that there are two fetal hearts in the measurement volume. The fitted two-element Gaussian model can then be used to determine which cluster, or in other words which fetal heart, the sample volume belongs to.

In a further example, clustering may be performed using a k-means clustering model.

In k-means clustering, there are no fundamental assumptions about how the positions of the sample volume are distributed. The k-means algorithm aims to divide the data into k subsets with minimal variance.

Various alternative clustering methods may be employed. For example, hierarchical clustering methods allow data to be grouped without specifying in advance the number of clusters to be generated. In the presence of maternal arteries within the sample volume, these clustering methods are useful and may be employed.

The determined depth of the fetal heartbeat and the positions of the at least two ultrasound transducers may then be used to locate the fetal heart.

FIG. 4 shows a schematic diagram of the ultrasound system described above applied to a maternal abdomen 180 containing twins. In other words, there are two fetal heartbeats within the imaging region.

As shown in FIG. 4, the ultrasound signal 250a produced by the first transducer 130a partially intersects both the first 260 and second 270 fetal hearts. Thus, echo signals received along these transmission lines will contain motion signals from both fetal hearts. This results in a large amount of noise in the received signal and reduces the accuracy of the determined location of the fetal heart.

Typically, this is addressed by providing multiple ultrasound probes, each directed to image one of the fetal hearts. It is difficult to place two separate ultrasound probes in the maternal abdomen for twin fetal heart rate monitoring.

For example, if both hearts are within the sample volume of one probe, the measured Doppler signal will reflect the motion of both hearts. Therefore, the Doppler signal exhibits multiple peaks and the algorithms used to estimate fetal heart rate, such as the autocorrelation function, are unable to determine the correct interbeat interval for each fetus.

Furthermore, since the position of both fetal hearts may change over time, this means that repositioning of the ultrasound transducers is required. Additionally, care needs to be taken to unambiguously assign the measured heart rate to the correct fetus. It is possible that the heart rate is incorrectly assigned. If the recorded heart rate trace looks suspicious, it may lead to the wrong intervention being chosen by the clinician.

If the measured heart rates happen to coincide or fall within the same range, the monitoring system may issue an alarm. Because it cannot determine whether it is really measuring both fetal heart rates or measuring the same fetal heart rate twice.

Ultrasound signals 250a respectively generated by the first transducer element 130a and the second transducer element 130b of the same transducer array 110 by employing a single flexible transducer array configured to conform to the subject's body. and 250b can illuminate both fetal hearts from different angles. Thus, it is possible to triangulate the position of both fetal hearts using at least one unshielded ultrasound transmission line. Thus, each fetal heart location can be accurately isolated from the remaining received echo signals, thereby increasing the overall accuracy of fetal heart region localization.

Since the transducer array can include any number of transducer elements each configured to transmit and receive ultrasound signals, at least one signal is transmitted by another Doppler source (such as another fetal heart). will not be interrupted.

Furthermore, instead of transmitting on all transducer elements simultaneously, only the group of elements directed toward the fetal heart (one or more elements with a given apodization profile) contributes to the total acoustic dose delivered to the fetus. May be activated for transmission to reduce.

Similarly, instead of being received by all elements simultaneously, only the group of elements (one or more elements with a given apodization profile) directed toward the fetal heart will have a signal-to-noise ratio of the received echo signal. can be activated for reception to improve the

FIG. 5 shows a 3D visualization of the sample volume of the measurement region containing two fetal hearts. FIG. 5 shows an array of ultrasound transducer elements 130 positioned over a measurement area defined by x, y and z axes. Within the measurement volume, a sample volume 280 is shown representing the division of the received echo signals with calculated Doppler power above a predetermined threshold. The darker the represented sample volume, the higher the Doppler power measured there, ie the more motion in this area.

The Doppler power measured across the transducer array can then be clustered and used to identify fetal heart regions 290 .

In other words, the Doppler power Pi,g is visualized in 3D, showing where the strongest Doppler signals are located. Visualization of Doppler power in 3D provides an easy way to see if there are only two Doppler sources in the measurement volume or if for example a pulsating maternal artery is present in the measurement volume. In particular, the maternal artery lies within the sample volume of the ultrasound transducer and can corrupt the acquired Doppler signal. This may result in erroneous registration of fetal heart rate, or alternatively maternal heart rate measurement. In this case, the system proposed above can identify the sources of such erroneous Doppler signals and discount them in the fetal heart signal.

In one example, all sample volume locations with Doppler power Pi,g greater than or equal to a predetermined threshold are fitted to a two-element Gaussian mixture model and then clustered. A median fetal heart rate is then calculated from the Doppler signals for each cluster to obtain two fetal heart rate measurements.

Such 3D visualization may be presented to the user via a suitable display, the display configured to show the fetal heart region to the user.

The example shown in FIG. 5 was obtained using a transducer array with 25 transducer elements. Doppler signals were calculated from 64 range gates. For experimental validation, the system was tested using an in-vitro setup of twin fetal hearts. The results confirm that a single flexible sensor matrix allows localization and visualization of twin fetal hearts. With the proposed clustering algorithm, two fetal heart rates can be detected without the aforementioned problem of manually positioning two ultrasound transducers. This can result in improved clinical workflow and better twin fetal health monitoring.

FIG. 6 shows an ultrasound imaging method 300 according to the invention.

At step 310, a plurality of ultrasound signals are acquired with at least two ultrasound transducers having different orientations with respect to the region of interest.

At step 320, the multiple ultrasound signals are split based on the depth of the signals using, for example, range gates.

At step 330, the Doppler power is calculated for each ultrasound signal split.

At step 340, the depth of the fetal heartbeat is determined for each ultrasound signal based on the Doppler power of the split of each ultrasound signal.

At step 350 the Doppler power is compared to a threshold power, and at step 360 Doppler power values equal to or greater than the threshold power may be clustered.

At step 370, a fetal heart region is identified based on the identified fetal heart rate and the locations of the at least two ultrasound transducers.

Variations to the disclosed embodiments can be understood and effected by those skilled in the art in practicing the claimed invention, from a study of the drawings, the disclosure, and the appended claims. In the claims, the word "comprising" does not exclude other elements or steps, and the indefinite articles "a" or "an" do not exclude a plurality. A single processor or other unit may fulfill the functions of several items recited in the claims. The mere fact that certain measures are recited in mutually different dependent claims does not indicate that a combination of these measures cannot be used to advantage. The computer program can be stored/distributed on any suitable medium, such as optical storage media or solid-state media provided with or as part of other hardware, but the Internet or other wired or wireless communications It can also be distributed in other forms, such as via a system. Any reference signs in the claims should not be construed as limiting the scope.

Claims (11)

An ultrasound system,
An ultrasound transducer array comprising a plurality of ultrasound transducer elements configured to conform to a body of a subject, wherein at least two ultrasound transducer elements of said plurality of ultrasound transducer elements are oriented in different directions in a region of interest. an ultrasonic transducer array that acquires a plurality of ultrasonic signals from
a processor that receives the plurality of ultrasonic signals acquired by the ultrasonic transducer array;
the processor
splitting the plurality of ultrasound signals based on signal depth;
Compute the Doppler power for each ultrasound signal split,
determining, for each ultrasound signal, the depth of the fetal heartbeat based on the Doppler power of the split of each ultrasound signal;
configured to identify a fetal heart region based on the determined fetal heart rate and the positions of the at least two ultrasound transducers;
determining the heartbeat depth of the fetus includes comparing the Doppler power to a threshold power;
determining the heartbeat depth of the fetus further comprises clustering the Doppler power above the threshold power ;
the clustering is performed using a two-element Gaussian mixture model or a k-means clustering model ;
ultrasound system. 2. The Doppler power calculation comprises calculating a Doppler signal based on the division of the ultrasound signal and calculating the Doppler power by calculating a root mean square value of the Doppler signal over a predetermined time period. An ultrasound system as described in . 3. The ultrasound system of claim 2, wherein said time period is one second or greater. 4. The ultrasound system of any of claims 1-3 , wherein the processor further calculates a median fetal heart rate based on the fetal heart rate of the identified fetal heart beats. 5. The ultrasound system of any of claims 1-4 , wherein the ultrasound system further comprises a display, the display showing the fetal heart region to the user. 6. The ultrasound of claim 5 , wherein the processor further determines the location of each ultrasound signal segment within the region of interest, and wherein the display shows the fetal heart region in relation to each ultrasound signal segment. sound wave system. 7. The ultrasound system of any of claims 1-6 , wherein the ultrasound transducer array further comprises a sensor, and wherein the processor determines the curvature of the ultrasound transducer array based on the output of the sensor. 8. The ultrasound system of Claim 7 , wherein the sensors comprise one or more of strain gauges, accelerometers, piezoelectric sensors and cameras. 9. The ultrasound system of any of claims 1-8 , wherein the plurality of ultrasound transducers comprises one or more of piezoelectric transducers and CMUTs. In the ultrasound imaging method,
acquiring a plurality of ultrasound signals using at least two ultrasound transducers with different orientations with respect to the region of interest;
dividing the plurality of ultrasound signals based on signal depth;
calculating the Doppler power for each ultrasound signal split;
determining, for each ultrasound signal, the depth of the fetal heartbeat based on the Doppler power of the split of each ultrasound signal;
comparing the Doppler power to a threshold power;
clustering the Doppler powers above the threshold power , wherein the clustering is performed using a two-element Gaussian mixture model or a k-means clustering model;
and determining a fetal heart region based on the determined fetal heart rate and the positions of the at least two ultrasound transducers. A computer program comprising computer program code means adapted to perform all the steps of the method according to claim 10 when run on a computer.

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